The present invention is generally concerned with high density pixel array systems, such as imagers, sensors, actuators, detectors and the like, and methods of fabricating such arrays, and is particularly concerned with a method of fabricating a high performance pixel array in exotic materials such as ferroelectric, piezoelectric, pyroelectric, acousto-optic materials and the like for integration in such a system.
In imaging systems for medical and other diagnostic sciences, an ongoing goal is the development of a low-cost, high quality, high resolution, real-time digital imaging system for an opaque target, such as the human body. Imaging systems are also used in non-medical applications such as nondestructive testing of materials and compounds. Such systems have the capability of providing on-line, non-invasive imaging. Currently, both X-ray and ultrasonic imaging techniques are used for displaying the internal characteristics of an opaque item, such as parts of human or animal bodies. Both techniques are subject to same disadvantages. In X-ray imaging, digitizing of an X-ray image directly is a challenge because silicon used in the pixels of focal plane digitizing arrays or detectors, such as charge coupled devices (CCDs), is damaged by X-rays. It is known to place a fluorescent or phosphorescent medium between the X-ray source and a visible detector matrix to convert the X-rays to visible light. There are still problems in using a screen of such material in front of the visible matrix detector. For example, if the screen is too thin, not enough of the X-rays will be absorbed and some will reach and damage the visible detector. If the screen is too thick, visible light is scattered, enlarging the area of the detector which is illuminated and reducing the digitized image resolution. In some cases, scattered light may escape without reaching the visible detect or at all. The fluorescent or phosphorescent material may also have non-uniform properties, degrading image quality and resolution. Some phosphorescent materials exhibit xe2x80x9cafter-glowxe2x80x9d, in other words they may continue to emit light even after the radiation source is no longer present. This may further degrade the image quality.
U.S. Pat. No. 5,519,227 of Karellas describes a structured scintillation screen which overcomes some of these problems. Regions of a transparent or semi-transparent scintillating substance are ablated to form an array of individual pixels. Each pixel is surrounded with an optically inactive material having a lower refractive index, so that the pixel is made to function as an optical waveguide. This confines the x-ray induced phosphorescence to t he individual pixels and channels it to the corresponding visible detector elements. This increases resolution and detection efficiency. The method of fabrication is as follows: The substrate of phosphorescent or optically active material is exposed to electromagnetic radiation, such as a laser beam, so as to ablate the substrate in exposed regions to produce a one or two dimensional array of pixels. A mask may be placed in contact with the substrate so that the desired regions are ablated by the laser beam. Following laser processing to form the pixels, the pixels are surrounded by an optically inactive interstitial material so as to avoid optical leakage from each pixel. The pixel structure is attached via a substrate to a visible detector such as a CCD camera.
t Other X-ray focal plane array (XFPA) medical matrix imagers have also been proposed, and have been introduced commercially in recent years, particularly for dental examinations. However, these imagers have, up to now, been very expensive and demonstrate marginal performance, due to the significant challenges in developing of a high performance, two dimensional XFPA detection matrix. One of the problems is that in order to replace high-resolution film radiography, the pixelated detector must have high uniformity and almost zero defects, with a resolution approaching 20 lp/mm, for good performance. All current commercial XFPA systems have demonstrated inferior imaging quality as compared with state-of-the-art commercial X-ray films, due to lack of sufficient resolution and low signal/noise.
A different imaging technique is ultrasound imaging, which has many applications. Such non-invasive ultrasonic imaging has advanced tremendously since its inception around 1950 and is currently one of the effective techniques for medical diagnostics of the internal human abdominal organs, the heart and great vessels. The transducer is the heart of all the medical ultrasound imagers. It performs the conversion of the electric signal into acoustic energy (transceiver) and, vice versa, translates back the received mechanical energy into an electric signal (receiver), to detect the information carried in the receiving signal. Consequently, there are fundamental relationships between the architecture and functional operation of the transducer and the quality of the resulting sonographic image.
Early transducers were based on a single element which was manually scanned. In the seventies, the linear phase-array transducers were introduced which were able to electronically focus and electronically steer the ultrasound beam in the plane of the linear array, by the application of suitable phase delays to each element. Current state-of-the-art clinical ultrasound imagers typically use linear phase-arrays (Nxc3x971) with more than N=100 elements, to electronically steer and focus the ultrasound beam. These ultrasound imagers are normally scanned in the B mode, which allows viewing of a cross-section slice. However, these arrays can only steer and focus in their elevation direction. Thus, in most cases, the lateral resolution in the azimuth direction can be completely different than in the elevation direction. This asymmetry of the ultrasound beam shape can make the detection of small cysts and lesions in the abdomen, fetus, or myocardium very difficult. In order to reduce the slice thickness and improve the elevation resolution, 1.5D phase-arrays, with (Nxc3x973, or (Nxc3x975) matrixes were recently implemented in certain ultrasound imaging systems. Currently, most advanced ultrasound imagers used for gynecology, obstetrics, encephalogy, opthamology and cardiology, are based on 1D, or 1.5D, electronic scanned B-mode phase-array transducers, which consist of one or a few rows of piezoelectric transducer elements, respectively. Full volume scanning is usually provided by mechanical scanning of the phase-array transducers either manually, or by a mechanical manipulator.
The most popular ultrasound body imaging is provided by the impulse-echo modality, in which the piezoelectric transducers act as both acoustic sources and detectors of the ultrasound radiation. The principle of the impulse-echo method is based on the ultrasonic transducer transmitting the sound impulse into the body (transceiver). The returning signal from the internal organ is detected by the transducer (receiver), which also determines the time lapse between the transmitted and received pulse, for the determination of the distance of the reflecting (scattering) organ. The further the organ from the origin, the longer is the time measured. Practically, in the impulse-echo method, the echo signals measure the changes in the reflected and scattered ultrasonic radiation, due to the acoustic-impedance differences at the borders between the various biological materials of the different tissues, to generate a mapping image, point by point. Consequently, the impulse-echo methodology is mostly utilized for non-invasive clinical imagery of soft internal tissues, which allows better penetration of these acoustic waves and their back return.
In the last 20 years, the image quality of medical ultrasound imaging has advanced sufficiently to make it an important, and sometimes indispensable, diagnostics modality in obstetrics and in the management of a large number of diseases. Nevertheless, current ultrasound imaging still suffers from a number of disadvantages, which are related to the marginal spatial resolution (blurred images), high noise components (noisy images) and manual operation (bulky scanners). This is also manifested in the subjectivity of the current ultrasound examinations, depending on the experience of the diagnostician to manipulate the ultrasound transducer and interpret the data. It is believed that, up to now, the full potential of ultrasound imagers has not yet been realized for real-time reliable clinical imaging for internal medicine. Nevertheless, although demonstrating only medium spatial resolution, with sometimes only fuzzy images, non-invasive ultrasound medical diagnostic modality is becoming more and more popular as compared with X-ray imaging. This is mostly due to the fact that this cost-effective ultrasound imaging modality does not involve ionizing radiation, while being safe and painless for the patient. The realization of the full potential of the ultrasound medical imaging modality will require a combination of advancements in the transducer architecture, together with improvements in its operating electronics design, its operating performance and in the computational post-processing (acquisition, reconstruction and rendering) of the data, to provide real-time, reliable, high-quality imaging of the targeted internal organs.
Currently, ultrasound is a very popular medical imaging modality, second only to conventional X-rays in the number of procedures performed. Its advantages over the other modalities, including conventional X-rays, computed tomography (CT) and magnetic resonance imaging (MRI), are that it is almost completely noninvasive, providing better soft-tissue differentiation than X-rays, capable of providing images in real-time, it is portable, and perhaps most important in todays environment of curtailing health-care costs, it is very cost-effective. However, even the most advanced ultrasound imaging devices today demonstrate only medium spatial resolution and limited signal-to-noise ratio to allow only a blurry view of internal human vessels. Additionally, current ultrasound imagers require mechanical maneuvering to obtain a volumetric 3D imager. In addition to medical imaging, ultrasound and other types of radiation are commonly used in therapeutic applications to the human or animal body. High intensity, focused ultrasound radiation beams are utilized to burn cancer cells, for example, by focusing the beam to the location of the malignancy.
At present, typical medical sonographic applications are related to the diagnostics in the following: internal medicine (liver, gall bladder and gall vessels, the pancreas, the spleen, the kidneys, the bladder and certain large blood vessels), gynecology, obstetrics, cardiology, opthamology and ultrasound guided aspiration. Typical non-medical sonographic applications are related to Nondestructive Evaluation (NDE), or Nondestructive Testing (NDT) of materials, especially avionic components, for detecting internal defects and/or materials fatigue.
A major step toward realizing improved ultrasound imaging and application is the engagement of two dimensional (2D) phase-array transducers for the impulse-echo operation, as described for example, in xe2x80x9cProgress in Two Dimensional Arrays for Real Time Volumetric Imagingxe2x80x9d by E. D. Light et al., Ultrasonic Imaging, 20, 1-15, 1998. Indeed, rectangular two dimensional phase-array transducers, which consist of a full matrix of pixels, are now emerging in RandD laboratories with a matrix of elements from 10xc3x9710 to 64xc3x9764. True 2D ultrasonic transducer phase-arrays are necessary for the following: improving the focus depth resolution, achieving completely electronic tuned high speed volumetric scanning and obtaining angle independent flow imaging without significant aberrations. Consequently, the introduction of the second dimension of the transducer allows elevation steering and dynamic focusing, as well as the use of phase aberration correction algorithms to reduce the B-scan slice thickness, thus achieving better volumetric imaging resolution. These 2D phase-array configurations must be carefully designed to achieve high medical imaging performance, while guaranteeing effective total manufacturability, including the array fabrication, the electronic integration and the compact packaging in a cost-effective manner.
Several 2D phase-array transducer architectures are currently being developed for medical diagnostics of critical lumens in the human body. For example, a square 2D transducer array operating at 5 MHz was developed to improve cardiac images, as the ultrasound beam can be electronically steered and dynamically focused to provide real-time three-dimensional monitoring. The electronic scanning of the ultrasound beam in a pyramidal pattern can display any desired plan sector, including the true short axis of the heart. Such medical transthorcic imaging of the heart has its challenges, as the acoustic window between the ribs limits the transducer footprint. This requires high-density arrays, due to the large number of miniature pixels necessary for shaping and transmitting the ultrasound beam and for the receiving of the echo. Additionally, due to the relatively large distance of the interesting location of the heart ( greater than 70 mm) from the external monitor, the attenuation of the body tissue plays a significant role, limiting the use of frequencies higher than 5 MHz as the absorption of the acoustic waves increases substantially at higher frequencies. This imposes certain restrictions on the maximum spatial resolution that the imager can resolve, due to the ultrasound waves diffraction laws. The intraluminal ultrasound device is an alternative to approach the desired lumen or area of the body more closely, thus avoiding the previous mentioned challenges. However, there are problems in introducing such a device into a human body, one of which is the very small dimensions required.
The manufacturing of very large numbers of such miniature transducer elements of the (Nxc3x97M) 2D phase-array matrix, concentrated in a small volume, provides an extreme challenge in engineering of the device and it""s manufacturing costs. Current reticulation techniques involve the use of high precision mechanical saws, as used in the semiconductor industry for dicing integrated circuit wafers. The blades of these saws are too wide to manufacture small pixels. Additionally, the high speed blades may cause cracks or chips in the pixels, and only low aspect ratio cuts can be achieved, since the saws cannot achieve sufficient groove depth to maintain the proper transducer modes of vibration. Chemical photo lithographic etching techniques are also generally unsuitable for etching of piezoelectric materials due to their inert chemical nature. Many other exotic materials like ferroelectrics, piezorestrictive, scintillator, phosphorous and fluorescent materials, which can serve as pixilated matrixes, will encounter the same challenges.
All these challenges have limited the current 2D phase-array transducers and therapeutic transmitters to matrixes with a small number of pixels, operating at low frequencies ( less than 2 MHz). For example, a conventional linear (1D) phase-array medical transducer, consisting of 64 operating elements, with the proper pixel size for 5 MHz operation, will require interelement spacing less than 0.15 mm. To make similar quality images in a fully-sampled 2D phase-array transducer, will require 64xc3x9764=4096 elements in an area of close to 9.6xc3x979.6 mm2. The fabrication of such a matrix of tiny elements with high aspect-radio dimensions is very difficult, and the current mechanical dicing Integrated Circuit technique to delineate common piezoelectric materials, fails to guarantee the integrity of the individual pixels.
It is an object of the present invention to provide a new and improved pixel array and method of fabricating such an array.
According to one aspect of the present invention, a method for micromilling a substrate to a predetermined depth is provided, which comprises the steps of directing a laser beam at a predetermined intensity at a surface of the substrate material, moving the substrate material relative to the laser beam at a constant velocity along a predetermined set of paths so as to remove the surface of the substrate material along the paths by application of a predetermined uniform flux per unit area, and repeating the relative movement of the substrate material and laser beam until the material has been ablated to the predetermined depth to form a series of grooves in a predetermined pattern, the grooves together defining a predetermined array of pixels separated by grooves.
The substrate material may be an X-ray fluorescent material selected from a group consisting of CdWO4, Bi1Ge3O12, YAG:Eu+3, YAG:Ce, CSI(Tl), CSI(Na), CSI, NaI, CsF, CaF(Eu), LiI(Eu), and Gd2SiO5Ce where an X-ray imaging detector is to be formed. In an alternative embodiment, the substrate material is an ultrasonic transducer material, such as a relaxor ferroelectric material, a piezoelectric or piezorestrictive, single crystal, or a piezoelectric ceramic. The relaxor ferroelectric single crystal material may be selected from the group consisting of (1-y) [Pb(Zn1/3Nb2/3)O3]+yPbTiO3 abbreviated PZN-PT, and (1-y) Pb(Mg1/3Nb2/3)O3+yPbTiO3 abbreviated PMN-PT, and other similar single crystals homologs, which demonstrate super high electromechanical coefficients (strain greater than 1%) and excellent coupling factors ( greater than 85%), such as those described in the article xe2x80x9cCan Relaxor Piezoelectric Materials Outperform PZT?xe2x80x9d, by Y. Yamashita et al., ISAF 1996, Proc 10th IEEE International Symposium on Applications of Ferroelectrics, East Brunswick, N.J., August 1996, pp. 71-77. Suitable single crystal piezoelectrics and piezorestrictors may be selected from the group consisting of Barium Titanate, BaTiO3, lithium tantalate, LiTaO3, and similar piezoelectric single crystals with high electromechanical coefficients. Suitable piezoelectric ceramics may have regular grain size (1 xcexcm to 25 xcexcm) or fine grain size, with good transducer characteristics. Suitable piezoelectric ceramic materials include materials based on different compositions of PZT (lead zirconate titanate, or (Pb[Zr,Ti]O3)), which are commercially available from various companies.
In an exemplary embodiment of the invention, the focus and intensity of the laser beam is varied as the depth of the ablated groove increases, in a manner to keep the energy introduced into the micro-milled material at a constant flux per unit area, or energy per unit time per unit area. This helps to ensure that a relatively smooth-sided groove is produced, and also that the adjacent material is not degraded or damaged by introduction of too much heat. The flux per unit area to be applied, in other words the power passing into the surface, is precisely determined depending on the melting point, opacity, and other critical properties of the material so that just enough energy is applied to ablate the material in the desired region without spreading outwardly from that region and potentially degrading adjacent pixels. The power or flux per unit area is controlled by means of the selected constant relative velocity, the laser beam intensity, and the laser focus. All of these may be adjusted to achieve the predetermined flux per unit area.
The grooves may be micro-milled in the substrate material in a rectangular x-y grid pattern, or in a curved or circular array pattern, forming a transducer array of cylindrical shape or round-cross-section. A circular array is particularly suitable for fitting into a catheter for approaching an organ within the body. The grooves separating the pixels are filled with glue material. The glue or adhesive material may be a material which is substantially reflective to visible light, in the case of an X-ray imaging array so as to optically isolate each pixel from adjacent pixels. The resultant pixelated substrate may be attached to a visible pixelated detector, such as a CCD (charge coupled device) detector or other visual matrix detector. In this way, x-rays incident on the pixel array of fluorescent or phosphorescent material will be converted to light rays, and these will be transmitted along each pixel to the underlying corresponding visible detecting pixel.
In the case of an ultrasonic transducer or transmitter array, the adhesive filling the grooves is of a flexible epoxy material. The pixel array may be covered with a matching layer for acoustic matching with the analyzed medium. Each pixel will have a suitable electrode secured at it""s lower end for connection of the pixels to supply (transceiver) and readout (receiver) electronics. Additional backing material may be provided to direct the acoustic energy toward the studied medium for energy effectiveness.
The laser micro-milling technique, in which the laser travels at constant velocity along each groove numerous times with a constant or controlled intensity to ensure uniform flux application per unit area gradually ablating the material to a greater and greater depth along each groove, ensures that the groove walls will be relatively smooth and uniform. Control of the laser focus and intensity helps to ensure smoothness of the walls. This technique enables micromilling of materials to form pixelated arrays of much smaller dimensions than was previously possible, without damaging the crystal material within the pixel due to laser thermal effects.
Preferably, the laser in the micro-milling process is controlled to be switched on only when the relative movement between the laser and target or substrate has reached a constant velocity. Additionally, the laser preferably has a pulsed output, and the first, large pulse (xe2x80x9cgiant pulsexe2x80x9d) when the laser is first switched on is xe2x80x9ckilledxe2x80x9d or removed in a conventional manner, to ensure that only uniform intensity laser pulses are impinged on the target. The relative velocity and pulse timing is such that adjacent laser pulses overlap to form a continuous groove along each desired straight or curved line or path.
According to another aspect of the present invention, an imaging array detector is provided, which comprises an array of micromilled, elongated pixels separated by grooves, an adhesive material filling the grooves between adjacent pixels, and each groove having a width in the range of 4 xcexcm to 25 xcexcm along its entire length.
In one embodiment of the invention, the detector is an X-ray focal plane imaging detector and the pixels are of an x-ray fluorescent material selected from the group consisting of CdWO4, Bi4Ge3O12, YAG:Eu+3, YAG:Ce, CSI(Tl), CSI(Na), CSI, NaI, CsF, CaF(Eu), LiI(Eu), and Gd2SiO5Ce. These materials were selected based on their laser micro-milling performance and X-ray detection performance. All of the materials listed above are found to fulfill a great part of the required criteria for effective scintillator material, and to be compatible with the laser micro-milling technique used to manufacture the reticulated array. Other materials with equivalent properties may alternatively be used in other embodiments of the invention.
In another embodiment of the invention, a transceiver and receiver is an ultrasonic imaging transducer, and the pixel phase array is formed from a suitable piezoelectric or piezorestrictive material, with the adhesive material filling the grooves being of a flexible glue material. The array may be rectangular, with straight grooves in x and y directions separating the pixels, or may be of circular or other shapes, with at least some of the grooves being curved
The method of this invention allows an array detector or transducer of very small dimensions but with a large number of pixel elements to be fabricated with high aspect ratio grooves having almost no material degradation. This results in a high density pixel array suitable for use as a detector, imager, sensor, beam transmitter, or the like, When used as an imager, the array produces high resolution imaging with excellent imaging quality.